Electrochemical Analyte Sensor

ABSTRACT

An electrochemical analyte sensor formed using conductive traces on a substrate can be used for determining and/or monitoring a level of analyte in in vitro or in vivo analyte-containing fluids. For example, an implantable sensor may be used for the continuous or automatic monitoring of a level of an analyte, such as glucose, lactate, or oxygen, in a patient. The electrochemical analyte sensor includes a substrate and conductive material disposed on the substrate, the conductive material forming a working electrode. In some sensors, the conductive material is disposed in recessed channels formed in a surface of the sensor. An electron transfer agent and/or catalyst may be provided to facilitate the electrolysis of the analyte or of a second compound whose level depends on the level of the analyte. A potential is formed between the working electrode and a reference electrode or counter/reference electrode and the resulting current is a function of the concentration of the analyte in the body fluid.

RELATED APPLICATIONS

This application is a continuation of U.S. patent application Ser. No.11/276,238 filed Feb. 20, 2006, which is a continuation of U.S. patentapplication Ser. No. 10/291,969 filed Nov. 11, 2002, now U.S. Pat. No.7,003,340, which is a continuation of U.S. patent application Ser. No.09/613,604 filed Jul. 10, 2000, now U.S. Pat. No. 6,484,046, which is acontinuation of U.S. patent application Ser. No. 09/034,372 filed Mar.4, 1998, now U.S. Pat. No. 6,134,461, the disclosures of each of whichare incorporated herein by reference for all purposes.

FIELD OF THE INVENTION

The present invention is, in general, directed to an analyte sensor.More particularly, the present invention relates to an electrochemicalsensor for determining a level of an analyte, such as glucose, lactate,or oxygen, in vivo and/or in vitro.

BACKGROUND OF THE INVENTION

The monitoring of the level of glucose or other analytes, such aslactate or oxygen, in certain individuals is vitally important to theirhealth. High or low levels of glucose or other analytes may havedetrimental effects. The monitoring of glucose is particularly importantto individuals with diabetes, as they must determine when insulin isneeded to reduce glucose levels in their bodies or when additionalglucose is needed to raise the level of glucose in their bodies.

A conventional technique used by many diabetics for personallymonitoring their blood glucose level includes the periodic drawing ofblood, the application of that blood to a test strip, and thedetermination of the blood glucose level using calorimetric,electrochemical, or photometric detection. This technique does notpermit continuous or automatic monitoring of glucose levels in the body,but typically must be performed manually on a periodic basis.Unfortunately, the consistency with which the level of glucose ischecked varies widely among individuals. Many diabetics find theperiodic testing inconvenient and they sometimes forget to test theirglucose level or do not have time for a proper test. In addition, someindividuals wish to avoid the pain associated with the test. Thesesituations may result in hyperglycemic or hypoglycemic episodes. An invivo glucose sensor that continuously or automatically monitors theindividual's glucose level would enable individuals to more easilymonitor their glucose, or other analyte, levels.

A variety of devices have been developed for continuous or automaticmonitoring of analytes, such as glucose, in the blood stream orinterstitial fluid. A number of these devices use electrochemicalsensors which are directly implanted into a blood vessel or in thesubcutaneous tissue of a patient. However, these devices are oftendifficult to reproducibly and inexpensively manufacture in largenumbers. In addition, these devices are typically large, bulky, and/orinflexible, and many can not be used effectively outside of a controlledmedical facility, such as a hospital or a doctor's office, unless thepatient is restricted in his activities.

The patient's comfort and the range of activities that can be performedwhile the sensor is implanted are important considerations in designingextended-use sensors for continuous or automatic in vivo monitoring ofthe level of an analyte, such as glucose. There is a need for a small,comfortable device which can continuously monitor the level of ananalyte, such as glucose, while still permitting the patient to engagein normal activities. Continuous and/or automatic monitoring of theanalyte can provide a warning to the patient when the level of theanalyte is at or near a threshold level. For example, if glucose is theanalyte, then the monitoring device might be configured to warn thepatient of current or impending hyperglycemia or hypoglycemia. Thepatient can then take appropriate actions.

In addition to in vivo monitoring of analyte levels, it is oftenimportant to determine the level of an analyte in a sample taken from asubject. For many individuals and for many analytes, continuousmonitoring of analyte level is not necessary, convenient, and/ordesirable. In vitro measurements are often useful in making periodicdeterminations of analyte level when an in vivo sensor is not beingused. Such measurements may also be useful for calibrating an in vivosensor. In these cases, it may be desirable to use small volume samplesdue to the difficulty of obtaining such samples, the discomfort of thepatient when the sample is obtained, and/or other reasons. However, mostconventional sensors are designed to test for analyte levels in sampleslarger than 3 microliters. It is desirable to have sensors that could beused for the in vitro monitoring of samples that may be as small as amicroliter, or even 25 nanoliters or less. The use of such small samplesreduces the inconvenience and pain associated with obtaining a sample,for example, by lancing a portion of the body to obtain a blood sample.

SUMMARY OF THE INVENTION

Generally, the present invention relates to an analyte sensor which canbe used for the in vivo and/or in vitro determination of a level of ananalyte in a fluid. Some embodiments of the invention are particularlyuseful for the continuous or automatic monitoring of a level of ananalyte, such as glucose or lactate, in a patient. One embodiment of theinvention is an electrochemical sensor. The electrochemical sensorincludes a substrate, a recessed channel formed in a surface of thesubstrate, and a conductive material disposed in the recessed channel.The conductive material forms a working electrode.

Another embodiment of the invention is an electrochemical sensor thatincludes a substrate and a plurality of recessed channels formed in atleast one surface of the substrate. Conductive material is disposed ineach of the recessed channels. The conductive material in at least oneof the recessed channels forms a working electrode.

A further embodiment of the invention is an analyte responsiveelectrochemical sensor that includes a working electrode and a masstransport limiting membrane. The mass transport limiting membranepreferably maintains a rate of permeation of the analyte through themass transport limiting membrane with a variation of less than 3% per °C. at temperatures ranging from 30° C. to 40° C.

Yet another embodiment of the invention is a method of determining alevel of an analyte in a fluid. The fluid is contacted by anelectrochemical sensor that includes a substrate, a recessed channel inthe substrate, and conductive material in the recessed channel forming aworking electrode. An electrical signal is generated by the sensor inresponse to the presence of the analyte. The level of the analyte may bedetermined from the electrical signal.

A further embodiment of the invention is a temperature sensor. Thetemperature sensor includes a substrate, a recessed channel formed inthe substrate, and a temperature probe disposed in the recessed channel.The temperature probe includes two probe leads that are disposed inspaced-apart portions of the recessed channel and atemperature-dependent element that is disposed in the recessed channeland is in contact with the two probe leads. The temperature-dependentelement is formed using a material having a temperature-dependentcharacteristic that alters a signal from the temperature probe inresponse to a change in temperature.

One embodiment of the invention is a method of determining a level of ananalyte in a fluid. The fluid is placed in contact with anelectrochemical sensor. The electrochemical sensor has a substrate, arecessed channel formed in a surface of the substrate, conductivematerial disposed in the recessed channel to form a working electrode,and a catalyst proximally disposed to the working electrode. A level ofa second compound in the fluid is changed by a reaction of the analytecatalyzed by the catalyst. A signal is generated in response to thelevel of the second electrode. The level of the analyte is determinedfrom the signal.

Another embodiment of the invention is an electrochemical sensor havinga substrate and a working electrode disposed on the substrate. Theworking electrode preferably includes a carbon material and has a widthalong at least a portion of the working electrode of 150 μm or less.

Another embodiment of the invention is an electrochemical sensor fordetermining a level of an analyte in a fluid. The electrochemical sensorincludes a substrate, a recessed channel formed in a surface of thesubstrate, and conductive material disposed in the recessed channel toform a working electrode. A catalyst is positioned near the workingelectrode to catalyze a reaction of the analyte which results in achange in a level of a second compound. The electrochemical sensorproduces a signal which is responsive to the level of the secondcompound.

Yet another embodiment of the invention is a sensor adapted forsubcutaneous implantation. The sensor includes a substrate, andconductive carbon non-leachably disposed on the substrate to form aworking electrode. An enzyme is non-leachably disposed in proximity tothe working electrode.

Another embodiment of the invention is an electrochemical sensorincluding a substrate and conductive material disposed on the substrate.The conductive material forms a plurality of traces. At least one of thetraces forms a working electrode. The plurality of conductive traces arepreferably separated on the surface of the substrate by a distance of150 μm or less.

One embodiment of the invention is an electrochemical sensor having asubstrate and conductive material disposed on a surface of thesubstrate. The conductive material forms a plurality of conductivetraces, at least one of which forms a working electrode. The pluralityof conductive traces are disposed on the surface of the substrate at apreferred density, along a width of the substrate, of 667 μm/trace orless.

Another embodiment of the invention is an electrochemical sensor havinga substrate, a conductive material disposed on the substrate to form aworking electrode, and a contact pad disposed on the substrate andoperatively connected to the working electrode. The contact pad is madeof a non-metallic conductive material to avoid or reduce corrosion.

Yet another embodiment of the invention is an analyte monitoring systemhaving a sensor and a control unit. The sensor includes a substrate, aworking electrode disposed on the substrate, and a contact pad coupledto the working electrode. The control unit has a conductive contactcoupled to the working electrode and is configured to apply a potentialacross the working electrode. At least one of the contact pad and theconductive contact is made using a non-metallic material to avoid orreduce corrosion.

A further embodiment of the invention is a method of determining a levelof an analyte in an animal. A sensor is implanted in the animal. Thesensor includes a substrate, a plurality of conductive traces disposedon the substrate, and a working electrode formed from one of theconductive traces. A signal is generated at the working electrode inresponse to the analyte. The level of the analyte is determined byanalyzing the signal. If the level of the analyte exceeds a thresholdamount, an electrical current is produced through a portion of theanimal to warn the animal. The electrical current is produced byapplying a potential between two of the conductive traces.

Another embodiment is an electrochemical sensor having a substrate, aconductive material disposed on the substrate to form a workingelectrode, and catalyst disposed in the conductive material. Thecatalyst catalyzes a reaction of the analyte to generate a signal at theworking electrode.

The above summary of the present invention is not intended to describeeach disclosed embodiment or every implementation of the presentinvention. The Figures and the detailed description which follow moreparticularly exemplify these embodiments.

BRIEF DESCRIPTION OF THE DRAWINGS

The invention may be more completely understood in consideration of thefollowing detailed description of various embodiments of the inventionin connection with the accompanying drawings, in which:

FIG. 1 is a block diagram of one embodiment of a subcutaneous analytemonitor using a subcutaneously implantable analyte sensor, according tothe invention;

FIG. 2 is a top view of one embodiment of an analyte sensor, accordingto the invention;

FIG. 3A is a cross-sectional view of the analyte sensor of FIG. 2;

FIG. 3B is a cross-sectional view of another embodiment of an analytesensor, according to the invention;

FIG. 4A is a cross-sectional view of a third embodiment of an analytesensor, according to the invention;

FIG. 4B is a cross-sectional view of a fourth embodiment of an analytesensor, according to the invention;

FIG. 5 is an expanded top view of a tip portion of the analyte sensor ofFIG. 2;

FIG. 6 is a cross-sectional view of a fifth embodiment of an analytesensor, according to the invention;

FIG. 7 is an expanded top view of a tip-portion of the analyte sensor ofFIG. 6;

FIG. 8 is an expanded bottom view of a tip-portion of the analyte sensorof FIG. 6;

FIG. 9 is a side view of the analyte sensor of FIG. 2;

FIG. 10 is a top view of the analyte sensor of FIG. 6;

FIG. 11 is a bottom view of the analyte sensor of FIG. 6;

FIG. 12 is another embodiment of an analyte sensor according to theinvention.

While the invention is amenable to various modifications and alternativeforms, specifics thereof have been shown by way of example in thedrawings and will be described in detail. It should be understood,however, that the intention is not to limit the invention to theparticular embodiments described. On the contrary, the intention is tocover all modifications, equivalents, and alternatives falling withinthe spirit and scope of the invention.

DETAILED DESCRIPTION OF THE INVENTION

The present invention is applicable to an analyte sensor for the in vivoand/or in vitro determination of an analyte, such as glucose, lactate,or oxygen, in a fluid. The analyte sensors of the present invention canbe utilized in a variety of contexts. For example, one embodiment of theanalyte sensor can be subcutaneously implanted in the interstitialtissue of a patient for the continuous or periodic monitoring of a levelof an analyte in a patient's interstitial fluid. This can then be usedto infer the analyte level in the patient's bloodstream. Other in vivoanalyte sensors can be made, according to the invention, for insertioninto an organ, vein, artery, or other portion of the body containingfluid. The in vivo analyte sensors may be configured for obtaining asingle measurement and/or for monitoring the level of the analyte over atime period which may range from hours to days or longer.

Another embodiment of the analyte sensor can be used for the in vitrodetermination of the presence and/or level of an analyte in a sample,and, particularly, in a small volume sample (e.g., 10 microliters to 50nanoliters or less). While the present invention is not so limited, anappreciation of various aspects of the invention may be gained through adiscussion of the examples provided below.

The following definitions are provided for terms used herein. A “counterelectrode” refers to an electrode paired with the working electrode,through which passes a current equal in magnitude and opposite in signto the current passing through the working electrode. In the context ofthe invention, the term “counter electrode” is meant to include counterelectrodes which also function as reference electrodes (i.e., acounter/reference electrode).

An “electrochemical sensor” is a device configured to detect thepresence and/or measure the level of an analyte in a sample viaelectrochemical oxidation and reduction reactions on the sensor. Thesereactions are transduced to an electrical signal that can be correlatedto an amount, concentration, or level of an analyte in the sample.

“Electrolysis” is the electrooxidation or electroreduction of a compoundeither directly at an electrode or via one or more electron transferagents.

A compound is “immobilized” on a surface when it is entrapped on orchemically bound to the surface.

A “non-leachable” or “non-releasable” compound or a compound that is“non-leachably disposed” is meant to define a compound that is affixedon the sensor such that it does not substantially diffuse away from theworking surface of the working electrode for the period in which thesensor is used (e.g., the period in which the sensor is implanted in apatient or measuring a sample).

Components are “immobilized” within a sensor, for example, when thecomponents are covalently, ionically, or coordinatively bound toconstituents of the sensor and/or are entrapped in a polymeric orsol-gel matrix or membrane which precludes mobility.

An “electron transfer agent” is a compound that carries electronsbetween the analyte and the working electrode, either directly, or incooperation with other electron transfer agents. One example of anelectron transfer agent is a redox mediator.

A “working electrode” is an electrode at which the analyte (or a secondcompound whose level depends on the level of the analyte) iselectrooxidized or electroreduced with or without the agency of anelectron transfer agent.

A “working surface” is that portion of the working electrode which iscoated with or is accessible to the electron transfer agent andconfigured for exposure to an analyte-containing fluid.

A “sensing layer” is a component of the sensor which includesconstituents that facilitate the electrolysis of the analyte. Thesensing layer may include constituents such as an electron transferagent, a catalyst which catalyzes a reaction of the analyte to produce aresponse at the electrode, or both. In some embodiments of the sensor,the sensing layer is non-leachably disposed in proximity to or on theworking electrode.

A “non-corroding” conductive material includes non-metallic materials,such as carbon and conductive polymers.

Analyte Sensor Systems

The sensors of the present invention can be utilized in a variety ofdevices and under a variety of conditions. The particular configurationof a sensor may depend on the use for which the sensor is intended andthe conditions under which the sensor will operate (e.g., in vivo or invitro). One embodiment of the analyte sensor is configured forimplantation into a patient or user for in vivo operation. For example,implantation of the sensor may be made in the arterial or venous systemsfor direct testing of analyte levels in blood. Alternatively, a sensormay be implanted in the interstitial tissue for determining the analytelevel in interstitial fluid. This level may be correlated and/orconverted to analyte levels in blood or other fluids. The site and depthof implantation may affect the particular shape, components, andconfiguration of the sensor. Subcutaneous implantation may be preferred,in some cases, to limit the depth of implantation of the sensor. Sensorsmay also be implanted in other regions of the body to determine analytelevels in other fluids.

An implantable analyte sensor may be used as part of an analytemonitoring system to continuously and/or periodically monitor the levelof an analyte in a body fluid of a patient. In addition to the sensor42, the analyte monitoring system 40 also typically includes a controlunit 44 for operating the sensor 42 (e.g., providing a potential to theelectrodes and obtaining measurements from the electrodes) and aprocessing unit 45 for analyzing the measurements from the sensor 42.The control unit 44 and processing unit 45 may be combined in a singleunit or may be separate.

Another embodiment of the sensor may be used for in vitro measurement ofa level of an analyte. The in vitro sensor is coupled to a control unitand/or a processing unit to form an analyte monitoring system. In someembodiments, an in vitro analyte monitoring system is also configured toprovide a sample to the sensor. For example, the analyte monitoringsystem may be configured to draw a sample from, for example, a lancedwound using a wicking and/or capillary action. The sample may then bedrawn into contact with the sensor. Examples of such sensors may befound in U.S. patent application Ser. No. 08/795,767 and PCT patentapplication Ser. No. PCT/US98/02652, incorporated herein by reference.

Other methods for providing a sample to the sensor include using a pump,syringe, or other mechanism to draw a sample from a patient throughtubing or the like either directly to the sensor or into a storage unitfrom which a sample is obtained for the sensor. The pump, syringe, orother mechanism may operate continuously, periodically, or when desiredto obtain a sample for testing. Other useful devices for providing ananalyte-containing fluid to the sensor include microfiltration and/ormicrodialysis devices. In some embodiments, particularly those using amicrodialysis device, the analyte may be drawn from the body fluidthrough a microporous membrane, for example, by osmotic pressure, into acarrier fluid which is then conveyed to the sensor for analysis. Otheruseful devices for acquiring a sample are those that collect body fluidstransported across the skin using techniques, such as reverseiontophoresis, to enhance the transport of fluid containing analyteacross the skin.

The Sensor

A sensor 42 includes at least one working electrode 58 formed on asubstrate 50, as shown in FIG. 2. The sensor 42 may also include atleast one counter electrode 60 (or counter/reference electrode) and/orat least one reference electrode 62 (see FIG. 8). The counter electrode60 and/or reference electrode 62 may be formed on the substrate 50 ormay be separate units. For example, the counter electrode and/orreference electrode may be formed on a second substrate which is alsoimplanted in the patient or, for some embodiments of the implantablesensors, the counter electrode and/or reference electrode may be placedon the skin of the patient with the working electrode or electrodesbeing implanted into the patient. The use of an on-the-skin counterand/or reference electrode with an implantable working electrode isdescribed in U.S. Pat. No. 5,593,852, incorporated herein by reference.

The working electrode or electrodes 58 are formed using conductivetraces 52 disposed on the substrate 50. The counter electrode 60 and/orreference electrode 62, as well as other optional portions of the sensor42, such as a temperature probe 66 (see FIG. 8), may also be formedusing conductive traces 52 disposed on the substrate 50. Theseconductive traces 52 may be formed over a smooth surface of thesubstrate 50 or within channels 54 formed by, for example, embossing,indenting or otherwise creating a depression in the substrate 50.

A sensing layer 64 (see FIGS. 3A and 3B) is often formed proximate to oron at least one of the working electrodes 58 to facilitate theelectrochemical detection of the analyte and the determination of itslevel in the sample fluid, particularly if the analyte can not beelectrolyzed at a desired rate and/or with a desired specificity on abare electrode. The sensing layer 64 may include an electron transferagent to transfer electrons directly or indirectly between the analyteand the working electrode 58. The sensing layer 64 may also contain acatalyst to catalyze a reaction of the analyte. The components of thesensing layer may be in a fluid or gel that is proximate to or incontact with the working electrode 58. Alternatively, the components ofthe sensing layer 64 may be disposed in a polymeric or sol-gel matrixthat is proximate to or on the working electrode 58. Preferably, thecomponents of the sensing layer 64 are non-leachably disposed within thesensor 42. More preferably, the components of the sensor 42 areimmobilized within the sensor 42.

In addition to the electrodes 58, 60, 62 and the sensing layer 64, thesensor 42 may also include a temperature probe 66 (see FIGS. 6 and 8), amass transport limiting layer 74 (see FIG. 9), a biocompatible layer 75(see FIG. 9), and/or other optional components, as described below. Eachof these items enhances the functioning of and/or results from thesensor 42, as discussed below.

The Substrate

The substrate 50 may be formed using a variety of non-conductingmaterials, including, for example, polymeric or plastic materials andceramic materials. Suitable materials for a particular sensor 42 may bedetermined, at least in part, based on the desired use of the sensor 42and properties of the materials.

In some embodiments, the substrate is flexible. For example, if thesensor 42 is configured for implantation into a patient, then the sensor42 may be made flexible (although rigid sensors may also be used forimplantable sensors) to reduce pain to the patient and damage to thetissue caused by the implantation of and/or the wearing of the sensor42. A flexible substrate 50 often increases the patient's comfort andallows a wider range of activities. Suitable materials for a flexiblesubstrate 50 include, for example, non-conducting plastic or polymericmaterials and other non-conducting, flexible, deformable materials.Examples of useful plastic or polymeric materials include thermoplasticssuch as polycarbonates, polyesters (e.g., Mylar™ and polyethyleneterephthalate (PET)), polyvinyl chloride (PVC), polyurethanes,polyethers, polyamides, polyimides, or copolymers of thesethermoplastics, such as PETG (glycol-modified polyethyleneterephthalate).

In other embodiments, the sensors 42 are made using a relatively rigidsubstrate 50 to, for example, provide structural support against bendingor breaking. Examples of rigid materials that may be used as thesubstrate 50 include poorly conducting ceramics, such as aluminum oxideand silicon dioxide. One advantage of an implantable sensor 42 having arigid substrate is that the sensor 42 may have a sharp point and/or asharp edge to aid in implantation of a sensor 42 without an additionalinsertion device. In addition, rigid substrates 50 may also be used insensors for in vitro analyte monitors.

It will be appreciated that for many sensors 42 and sensor applications,both rigid and flexible sensors will operate adequately. The flexibilityof the sensor 42 may also be controlled and varied along a continuum bychanging, for example, the composition and/or thickness of the substrate50.

In addition to considerations regarding flexibility, it is oftendesirable that implantable sensors 42 should have a substrate 50 whichis non-toxic. Preferably, the substrate 50 is approved by one or moreappropriate governmental agencies or private groups for in vivo use.

The sensor 42 may include optional features to facilitate insertion ofan implantable sensor 42, as shown in FIG. 12. For example, the sensor42 may be pointed at the tip 123 to ease insertion. In addition, thesensor 42 may include a barb 125 which assists in anchoring the sensor42 within the tissue of the patient during operation of the sensor 42.However, the barb 125 is typically small enough that little damage iscaused to the subcutaneous tissue when the sensor 42 is removed forreplacement.

Although the substrate 50 in at least some embodiments has uniformdimensions along the entire length of the sensor 42, in otherembodiments, the substrate 50 has a distal end 67 and a proximal end 65with different widths 53, 55, respectively, as illustrated in FIG. 2. Inthese embodiments, the distal end 67 of the substrate 50 may have arelatively narrow width 53. For sensors 42 which are implantable intothe subcutaneous tissue or another portion of a patient's body, thenarrow width 53 of the distal end 67 of the substrate 50 may facilitatethe implantation of the sensor 42. Often, the narrower the width of thesensor 42, the less pain the patient will feel during implantation ofthe sensor and afterwards.

For subcutaneously implantable sensors 42 which are designed forcontinuous or periodic monitoring of the analyte during normalactivities of the patient, a distal end 67 of the sensor 42 which is tobe implanted into the patient has a width 53 of 2 mm or less, preferably1 mm or less, and more preferably 0.5 mm or less. If the sensor 42 doesnot have regions of different widths, then the sensor 42 will typicallyhave an overall width of, for example, 2 mm, 1.5 mm, 1 mm, 0.5 mm, 0.25mm, or less. However, wider or narrower sensors may be used. Inparticular, wider implantable sensors may be used for insertion intoveins or arteries or when the movement of the patient is limited, forexample, when the patient is confined in bed or in a hospital.

For sensors 42 which are designed for measuring small volume in vitrosamples, the narrow width 53 may reduce the volume of sample needed foran accurate reading. The narrow width 53 of the sensor 42 results in allof the electrodes of the sensor 42 being closely congregated, therebyrequiring less sample volume to cover all of the electrodes. The widthof an in vitro sensor 42 may vary depending, at least in part, on thevolume of sample available to the sensor 42 and the dimensions of thesample chamber in which the sensor 42 is disposed.

Returning to FIG. 2, the proximal end 65 of the sensor 42 may have awidth 55 larger than the distal end 67 to facilitate the connectionbetween contact pads 49 of the electrodes and contacts on a controlunit. The wider the sensor 42 at this point, the larger the contact pads49 can be made. This may reduce the precision needed to properly connectthe sensor 42 to contacts on the control unit (e.g., sensor control unit44 of FIG. 1). However, the maximum width of the sensor 42 may beconstrained so that the sensor 42 remains small for the convenience andcomfort of the patient and/or to fit the desired size of the analytemonitor. For example, the proximal end 65 of a subcutaneouslyimplantable sensor 42, such as the sensor 42 illustrated in FIG. 1, mayhave a width 55 ranging from 0.5 mm to 15 mm, preferably from 1 mm to 10mm, and more preferably from 3 mm to 7 mm. However, wider or narrowersensors may be used in this and other in vivo and in vitro applications.

The thickness of the substrate 50 may be determined by the mechanicalproperties of the substrate material (e.g., the strength, modulus,and/or flexibility of the material), the desired use of the sensor 42including stresses on the substrate 50 arising from that use, as well asthe depth of any channels or indentations formed in the substrate 50, asdiscussed below. Typically, the substrate 50 of a subcutaneouslyimplantable sensor 42 for continuous or periodic monitoring of the levelof an analyte while the patient engages in normal activities has athickness of 50 to 500 μm and preferably 100 to 300 μm. However, thickerand thinner substrates 50 may be used, particularly in other types of invivo sensors 42.

The length of the sensor 42 may have a wide range of values depending ona variety of factors. Factors which influence the length of animplantable sensor 42 may include the depth of implantation into thepatient and the ability of the patient to manipulate a small flexiblesensor 42 and make connections between the sensor 42 and the sensorcontrol unit 44. A subcutaneously implantable sensor 42 for the analytemonitor illustrated in FIG. 1 may have a length ranging from 0.3 to 5cm, however, longer or shorter sensors may be used. The length of thenarrow portion of the sensor 42 (e.g., the portion which issubcutaneously inserted into the patient), if the sensor 42 has narrowand wide portions, is typically about 0.25 to 2 cm in length. However,longer and shorter portions may be used. All or only a part of thisnarrow portion may be subcutaneously implanted into the patient.

The lengths of other implantable sensors 42 will vary depending, atleast in part, on the portion of the patient into which the sensor 42 isto be implanted or inserted. The length of in vitro sensors may varyover a wide range depending on the particular configuration of theanalyte monitoring system and, in particular, the distance between thecontacts of the control unit and the sample.

Conductive Traces

At least one conductive trace 52 is formed on the substrate for use inconstructing a working electrode 58. In addition, other conductivetraces 52 may be formed on the substrate 50 for use as electrodes (e.g.,additional working electrodes, as well as counter, counter/reference,and/or reference electrodes) and other components, such as a temperatureprobe. The conductive traces 52 may extend most of the distance along alength 57 of the sensor 50, as illustrated in FIG. 2, although this isnot necessary. The placement of the conductive traces 52 may depend onthe particular configuration of the analyte monitoring system (e.g., theplacement of control unit contacts and/or the sample chamber in relationto the sensor 42). For implantable sensors, particularly subcutaneouslyimplantable sensors, the conductive traces typically extend close to thetip of the sensor 42 to minimize the amount of the sensor that must beimplanted.

The conductive traces 52 may be formed on the substrate 50 by a varietyof techniques, including, for example, photolithography, screenprinting, or other impact or non-impact printing techniques. Theconductive traces 52 may also be formed by carbonizing conductive traces52 in an organic (e.g., polymeric or plastic) substrate 50 using alaser. A description of some exemplary methods for forming the sensor 42is provided in U.S. Pat. No. 6,103,033, incorporated herein byreference.

Another method for disposing the conductive traces 52 on the substrate50 includes the formation of recessed channels 54 in one or moresurfaces of the substrate 50 and the subsequent filling of theserecessed channels 54 with a conductive material 56, as shown in FIG. 3A.The recessed channels 54 may be formed by indenting, embossing, orotherwise creating a depression in the surface of the substrate 50.Exemplary methods for forming channels and electrodes in a surface of asubstrate can be found in U.S. Pat. No. 6,103,033. The depth of thechannels is typically related to the thickness of the substrate 50. Inone embodiment, the channels have depths in the range of about 12.5 to75 μm (0.5 to 3 mils), and preferably about 25 to 50 μm (1 to 2 mils).

The conductive traces are typically formed using a conductive material56 such as carbon (e.g., graphite), a conductive polymer, a metal oralloy (e.g., gold or gold alloy), or a metallic compound (e.g.,ruthenium dioxide or titanium dioxide). The formation of films ofcarbon, conductive polymer, metal, alloy, or metallic compound arewell-known and include, for example, chemical vapor deposition (CVD),physical vapor deposition, sputtering, reactive sputtering, printing,coating, and painting. The conductive material 56 which fills thechannels 54 is often formed using a precursor material, such as aconductive ink or paste. In these embodiments, the conductive material56 is deposited on the substrate 50 using methods such as coating,painting, or applying the material using a spreading instrument, such asa coating blade. Excess conductive material between the channels 54 isthen removed by, for example, running a blade along the substratesurface.

In one embodiment, the conductive material 56 is a part of a precursormaterial, such as a conductive ink, obtainable, for example, from Ercon,Inc. (Wareham, Mass.), Metech, Inc. (Elverson, Pa.), E.I. du Pont deNemours and Co. (Wilmington, Del.), Emca-Remex Products(Montgomeryville, Pa.), or MCA Services (Melbourn, Great Britain). Theconductive ink is typically applied as a semiliquid or paste whichcontains particles of the carbon, metal, alloy, or metallic compound anda solvent or dispersant. After application of the conductive ink on thesubstrate 50 (e.g., in the channels 54), the solvent or dispersantevaporates to leave behind a solid mass of conductive material 56.

In addition to the particles of carbon, metal, alloy, or metalliccompound, the conductive ink may also contain a binder. The binder mayoptionally be cured to further bind the conductive material 56 withinthe channel 54 and/or on the substrate 50. Curing the binder increasesthe conductivity of the conductive material 56. However, this istypically not necessary as the currents carried by the conductivematerial 56 within the conductive traces 52 are often relatively low(usually less than 1 μA and often less than 100 nA). Typical bindersinclude, for example, polyurethane resins, cellulose derivatives,elastomers, and highly fluorinated polymers. Examples of elastomersinclude silicones, polymeric dienes, and acrylonitrile-butadiene-styrene(ABS) resins. One example of a fluorinated polymer binder is Teflon®(DuPont, Wilmington, Del.). These binders are cured using, for example,heat or light, including ultraviolet (UV) light. The appropriate curingmethod typically depends on the particular binder which is used.

Often, when a liquid or semiliquid precursor of the conductive material56 (e.g., a conductive ink) is deposited in the channel 54, theprecursor fills the channel 54. However, when the solvent or dispersantevaporates, the conductive material 56 which remains may lose volumesuch that the conductive material 56 may or may not continue to fill thechannel 54. Preferred conductive materials 56 do not pull away from thesubstrate 50 as they lose volume, but rather decrease in height withinthe channel 54. These conductive materials 56 typically adhere well tothe substrate 50 and therefore do not pull away from the substrate 50during evaporation of the solvent or dispersant. Other suitableconductive materials 56 either adhere to at least a portion of thesubstrate 50 and/or contain another additive, such as a binder, whichadheres the conductive material 56 to the substrate 50. Preferably, theconductive material 56 in the channels 54 is non-leachable, and morepreferably immobilized on the substrate 50. In some embodiments, theconductive material 56 may be formed by multiple applications of aliquid or semiliquid precursor interspersed with removal of the solventor dispersant.

In another embodiment, the channels 54 are formed using a laser. Thelaser carbonizes the polymer or plastic material. The carbon formed inthis process is used as the conductive material 56. Additionalconductive material 56, such as a conductive carbon ink, may be used tosupplement the carbon formed by the laser.

In a further embodiment, the conductive traces 52 are formed by padprinting techniques. For example, a film of conductive material isformed either as a continuous film or as a coating layer deposited on acarrier film. This film of conductive material is brought between aprint head and the substrate 50. A pattern on the surface of thesubstrate 50 is made using the print head according to a desired patternof conductive traces 52. The conductive material is transferred bypressure and/or heat from the film of conductive material to thesubstrate 50. This technique often produces channels (e.g., depressionscaused by the print head) in the substrate 50. Alternatively, theconductive material is deposited on the surface of the substrate 50without forming substantial depressions.

In other embodiments, the conductive traces 52 are formed by non-impactprinting techniques. Such techniques include electrophotography andmagnetography. In these processes, an image of the conductive traces 52is electrically or magnetically formed on a drum. A laser or LED may beused to electrically form an image. A magnetic recording head may beused to magnetically form an image. A toner material (e.g., a conductivematerial, such as a conductive ink) is then attracted to portions of thedrum according to the image. The toner material is then applied to thesubstrate by contact between the drum and the substrate. For example,the substrate may be rolled over the drum. The toner material may thenbe dried and/or a binder in the toner material may be cured to adherethe toner material to the substrate.

Another non-impact printing technique includes ejecting droplets ofconductive material onto the substrate in a desired pattern. Examples ofthis technique include ink jet printing and piezo jet printing. An imageis sent to the printer which then ejects the conductive material (e.g.,a conductive ink) according to the pattern. The printer may provide acontinuous stream of conductive material or the printer may eject theconductive material in discrete amounts at the desired points.

Yet another non-impact printing embodiment of forming the conductivetraces includes an ionographic process. In the this process, a curable,liquid precursor, such as a photopolymerizable acrylic resin (e.g.,Solimer 7501 from Cubital, Bad Kreuznach, Germany) is deposited over asurface of a substrate 50. A photomask having a positive or negativeimage of the conductive traces 52 is then used to cure the liquidprecursor. Light (e.g., visible or ultraviolet light) is directedthrough the photomask to cure the liquid precursor and form a solidlayer over the substrate according to the image on the photomask.Uncured liquid precursor is removed leaving behind channels 54 in thesolid layer. These channels 54 can then be filled with conductivematerial 56 to form conductive traces 52.

Conductive traces 52 (and channels 54, if used) can be formed withrelatively narrow widths, for example, in the range of 25 to 250 μm, andincluding widths of, for example, 250 μm, 150 μm, 100 μm, 75 μm, 50 μm,25 μm or less by the methods described above. In embodiments with two ormore conductive traces 52 on the same side of the substrate 50, theconductive traces 52 are separated by distances sufficient to preventconduction between the conductive traces 52. The edge-to-edge distancebetween the conductive traces is preferably in the range of 25 to 250 μmand may be, for example, 150 μm, 100 μm, 75 μm, 50 μm, or less. Thedensity of the conductive traces 52 on the substrate 50 is preferably inthe range of about 150 to 700 μm/trace and may be as small as 667μm/trace or less, 333 μm/trace or less, or even 167 μm/trace or less.

The working electrode 58 and the counter electrode 60 (if a separatereference electrode is used) are often made using a conductive material56, such as carbon. Suitable carbon conductive inks are available fromErcon, Inc. (Wareham, Mass.), Metech, Inc. (Elverson, Pa.), E.I. du Pontde Nemours and Co. (Wilmington, Del.), Emca-Remex Products(Montgomeryville, Pa.), or MCA Services (Melbourn, Great Britain).Typically, the working surface 51 of the working electrode 58 is atleast a portion of the conductive trace 52 that is in contact with theanalyte-containing fluid (e.g., implanted in the patient).

The reference electrode 62 and/or counter/reference electrode aretypically formed using conductive material 56 that is a suitablereference material, for example silver/silver chloride or anon-leachable redox couple bound to a conductive material, for example,a carbon-bound redox couple. Suitable silver/silver chloride conductiveinks are available from Ercon, Inc. (Wareham, Mass.), Metech, Inc.(Elverson, Pa.), E.I. du Pont de Nemours and Co. (Wilmington, Del.),Emca-Remex Products (Montgomeryville, Pa.), or MCA Services (Melbourn,Great Britain). Silver/silver chloride electrodes illustrate a type ofreference electrode that involves the reaction of a metal electrode witha constituent of the sample or body fluid, in this case, Cl⁻.

Suitable redox couples for binding to the conductive material of thereference electrode include, for example, redox polymers (e.g., polymershaving multiple redox centers.) It is preferred that the referenceelectrode surface be non-corroding so that an erroneous potential is notmeasured. Preferred conductive materials include less corrosive metals,such as gold and palladium. Most preferred are non-corrosive materialsincluding non-metallic conductors, such as carbon and conductingpolymers. A redox polymer can be adsorbed on or covalently bound to theconductive material of the reference electrode, such as a carbon surfaceof a conductive trace 52. Non-polymeric redox couples can be similarlybound to carbon or gold surfaces.

A variety of methods may be used to immobilize a redox polymer on anelectrode surface. One method is adsorptive immobilization. This methodis particularly useful for redox polymers with relatively high molecularweights. The molecular weight of a polymer may be increased, forexample, by cross-linking.

Another method for immobilizing the redox polymer includes thefunctionalization of the electrode surface and then the chemicalbonding, often covalently, of the redox polymer to the functional groupson the electrode surface. One example of this type of immobilizationbegins with a poly(4-vinylpyridine). The polymer's pyridine rings are,in part, complexed with a reducible/oxidizable species, such as[Os(bpy)₂Cl]^(+/2+) where bpy is 2,2′-bipyridine. Part of the pyridinerings are quaternized by reaction with 2-bromoethylamine. The polymer isthen crosslinked, for example, using a diepoxide, such as polyethyleneglycol diglycidyl ether.

Carbon surfaces can be modified for attachment of a redox species orpolymer, for example, by electroreduction of a diazonium salt. As anillustration, reduction of a diazonium salt formed upon diazotization ofp-aminobenzoic acid modifies a carbon surface with phenylcarboxylic acidfunctional groups. These functional groups can then be activated by acarbodiimide, such as 1-ethyl-3-(3-dimethylaminopropyl)-carbodiimidehydrochloride. The activated functional groups are then bound with aamine-functionalized redox couple, such as the quaternizedosmium-containing redox polymer described above or2-aminoethylferrocene, to form the redox couple.

Similarly, gold can be functionalized by an amine, such as cystamine. Aredox couple such as [Os(bpy)₂(pyridine-4-carboxylate)Cl]^(0/+) isactivated by 1-ethyl-3-(3-dimethylaminopropyl)-carbodiimidehydrochloride to form a reactive O-acylisourea which reacts with thegold-bound amine to form an amide.

In one embodiment, in addition to using the conductive traces 52 aselectrodes or probe leads, two or more of the conductive traces 52 onthe substrate 50 are used to give the patient a mild electrical shockwhen, for example, the analyte level exceeds a threshold level. Thisshock may act as a warning or alarm to the patient to initiate someaction to restore the appropriate level of the analyte.

The mild electrical shock is produced by applying a potential betweenany two conductive traces 52 that are not otherwise connected by aconductive path. For example, two of the electrodes 58, 60, 62 or oneelectrode 58, 60, 62 and the temperature probe 66 may be used to providethe mild shock. Preferably, the working electrode 58 and the referenceelectrode 62 are not used for this purpose as this may cause some damageto the chemical components on or proximate to the particular electrode(e.g., the sensing layer on the working electrode or the redox couple onthe reference electrode).

The current used to produce the mild shock is typically 0.1 to 1 mA.Higher or lower currents may be used, although care should be taken toavoid harm to the patient. The potential between the conductive tracesis typically 1 to 10 volts. However, higher or lower voltages may beused depending, for example, on the resistance of the conductive traces52, the distance between the conductive traces 52 and the desired amountof current. When the mild shock is delivered, potentials at the workingelectrode 58 and across the temperature probe 66 may be removed toprevent harm to those components caused by unwanted conduction betweenthe working electrode 58 (and/or temperature probe 66, if used) and theconductive traces 52 which provide the mild shock.

Contact Pads

Typically, each of the conductive traces 52 includes a contact pad 49.The contact pad 49 may simply be a portion of the conductive trace 52that is indistinguishable from the rest of the trace 52 except that thecontact pad 49 is brought into contact with the conductive contacts of acontrol unit (e.g., the sensor control unit 44 of FIG. 1). Morecommonly, however, the contact pad 49 is a region of the conductivetrace 52 that has a larger width than other regions of the trace 52 tofacilitate a connection with the contacts on the control unit. By makingthe contact pads 49 relatively large as compared with the width of theconductive traces 52, the need for precise registration between thecontact pads 49 and the contacts on the control unit is less criticalthan with small contact pads.

The contact pads 49 are typically made using the same material as theconductive material 56 of the conductive traces 52. However, this is notnecessary. Although metal, alloys, and metallic compounds may be used toform the contact pads 49, in some embodiments, it is desirable to makethe contact pads 49 from a carbon or other non-metallic material, suchas a conducting polymer. In contrast to metal or alloy contact pads,carbon and other non-metallic contact pads are not easily corroded ifthe contact pads 49 are in a wet, moist, or humid environment. Metalsand alloys may corrode under these conditions, particularly if thecontact pads 49 and contacts of the control unit are made usingdifferent metals or alloys. However, carbon and non-metallic contactpads 49 do not significantly corrode, even if the contacts of thecontrol device are metal or alloy.

One embodiment of the invention includes a sensor 42 having contact pads49 and a control unit 44 having conductive contacts (not shown). Duringoperation of the sensor 42, the contact pads 49 and conductive contactsare in contact with each other. In this embodiment, either the contactpads 49 or the conductive contacts are made using a non-corroding,conductive material. Such materials include, for example, carbon andconducting polymers. Preferred non-corroding materials include graphiteand vitreous carbon. The opposing contact pad or conductive contact ismade using carbon, a conducting polymer, a metal, such as gold,palladium, or platinum group metal, or a metallic compound, such asruthenium dioxide. This configuration of contact pads and conductivecontacts typically reduces corrosion. Preferably, when the sensor isplaced in a 3 mM, and more preferably, in a 100 mM, NaCl solution, thesignal arising due to the corrosion of the contact pads and/orconductive contacts is less than 3% of the signal generated by thesensor when exposed to concentration of analyte in the normalphysiological range. For at least some subcutaneous glucose sensors, thecurrent generated by analyte in a normal physiological range ranges from3 to 500 nA.

Each of the electrodes 58, 60, 62, as well as the two probe leads 68, 70of the temperature probe 66 (described below), are connected to contactpads 49 as shown in FIGS. 10 and 11. In one embodiment (not shown), thecontact pads 49 are on the same side of the substrate 50 as therespective electrodes or temperature probe leads to which the contactpads 49 are attached.

In other embodiments, the conductive traces 52 on at least one side areconnected through vias in the substrate to contact pads 49 a on theopposite surface of the substrate 50, as shown in FIGS. 10 and 11. Anadvantage of this configuration is that contact between the contacts onthe control unit and each of the electrodes 58, 60, 62 and the probeleads 68,70 of the temperature probe 66 can be made from a single sideof the substrate 50.

In yet other embodiments (not shown), vias through the substrate areused to provide contact pads on both sides of the substrate 50 for eachconductive trace 52. The vias connecting the conductive traces 52 withthe contact pads 49 a can be formed by making holes through thesubstrate 50 at the appropriate points and then filling the holes withconductive material 56.

Exemplary Electrode Configurations

A number of exemplary electrode configurations are described below,however, it will be understood that other configurations may also beused. In one embodiment, illustrated in FIG. 3A, the sensor 42 includestwo working electrodes 58 a, 58 b and one counter electrode 60, whichalso functions as a reference electrode. In another embodiment, thesensor includes one working electrode 58 a, one counter electrode 60,and one reference electrode 62, as shown in FIG. 3B. Each of theseembodiments is illustrated with all of the electrodes formed on the sameside of the substrate 50.

Alternatively, one or more of the electrodes may be formed on anopposing side of the substrate 50. This may be convenient if theelectrodes are formed using two different types of conductive material56 (e.g., carbon and silver/silver chloride). Then, at least in someembodiments, only one type of conductive material 56 needs to be appliedto each side of the substrate 50, thereby reducing the number of stepsin the manufacturing process and/or easing the registration constraintsin the process. For example, if the working electrode 58 is formed usinga carbon-based conductive material 56 and the reference orcounter/reference electrode is formed using a silver/silver chlorideconductive material 56, then the working electrode and reference orcounter/reference electrode may be formed on opposing sides of thesubstrate 50 for case of manufacture.

In another embodiment, two working electrodes 58 and one counterelectrode 60 are formed on one side of the substrate 50 and onereference electrode 62 and a temperature probe 66 are formed on anopposing side of the substrate 50, as illustrated in FIG. 6. Theopposing sides of the tip of this embodiment of the sensor 42 areillustrated in FIGS. 7 and 8.

Sensing Layer

Some analytes, such as oxygen, can be directly electrooxidized orelectroreduced on the working electrode 58. Other analytes, such asglucose and lactate, require the presence of at least one electrontransfer agent and/or at least one catalyst to facilitate theelectrooxidation or electroreduction of the analyte. Catalysts may alsobe used for those analyte, such as oxygen, that can be directlyelectrooxidized or electroreduced on the working electrode 58. For theseanalytes, each working electrode 58 has a sensing layer 64 formedproximate to or on a working surface of the working electrode 58.Typically, the sensing layer 64 is formed near or on only a smallportion of the working electrode 58, often near a tip of the sensor 42.This limits the amount of material needed to form the sensor 42 andplaces the sensing layer 64 in the best position for contact with theanalyte-containing fluid (e.g., a body fluid, sample fluid, or carrierfluid).

The sensing layer 64 includes one or more components designed tofacilitate the electrolysis of the analyte. The sensing layer 64 mayinclude, for example, a catalyst to catalyze a reaction of the analyteand produce a response at the working electrode 58, an electron transferagent to indirectly or directly transfer electrons between the analyteand the working electrode 58, or both.

The sensing layer 64 may be formed as a solid composition of the desiredcomponents (e.g., an electron transfer agent and/or a catalyst). Thesecomponents are preferably non-leachable from the sensor 42 and morepreferably are immobilized on the sensor 42. For example, the componentsmay be immobilized on a working electrode 58. Alternatively, thecomponents of the sensing layer 64 may be immobilized within or betweenone or more membranes or films disposed over the working electrode 58 orthe components may be immobilized in a polymeric or sol-gel matrix.

Examples of immobilized sensing layers are described in U.S. Pat. Nos.5,262,035, 5,264,104, 5,264,105, 5,320,725, 5,593,852, and 5,665,222,and PCT Patent Application No. US98/02403 entitled “Soybean PeroxidaseElectrochemical Sensor”, filed on Feb. 11, 1998, published asWO-1998/035053, incorporated herein by reference.

In some embodiments, one or more of the components of the sensing layer64 may be solvated, dispersed, or suspended in a fluid within thesensing layer 64, instead of forming a solid composition. The fluid maybe provided with the sensor 42 or may be absorbed by the sensor 42 fromthe analyte-containing fluid. Preferably, the components which aresolvated, dispersed, or suspended in this type of sensing layer 64 arenon-leachable from the sensing layer. Non-leachability may beaccomplished, for example, by providing barriers (e.g., the electrode,substrate, membranes, and/or films) around the sensing layer whichprevent the leaching of the components of the sensing layer 64. Oneexample of such a barrier is a microporous membrane or film which allowsdiffusion of the analyte into the sensing layer 64 to make contact withthe components of the sensing layer 64, but reduces or eliminates thediffusion of the sensing layer components (e.g., a electron transferagent and/or a catalyst) out of the sensing layer 64.

A variety of different sensing layer configurations can be used. In oneembodiment, the sensing layer 64 is deposited on the conductive material56 of a working electrode 58 a, as illustrated in FIGS. 3A and 3B. Thesensing layer 64 may extend beyond the conductive material 56 of theworking electrode 58 a. In some cases, the sensing layer 64 may alsoextend over the counter electrode 60 or reference electrode 62 withoutdegrading the performance of the glucose sensor. For those sensors 42which utilize channels 54 within which the conductive material 56 isdeposited, a portion of the sensing layer 64 may be formed within thechannel 54 if the conductive material 56 does not fill the channel 54.

A sensing layer 64 in direct contact with the working electrode 58 a maycontain an electron transfer agent to transfer electrons directly orindirectly between the analyte and the working electrode, as well as acatalyst to facilitate a reaction of the analyte. For example, aglucose, lactate, or oxygen electrode may be formed having a sensinglayer which contains a catalyst, such as glucose oxidase, lactateoxidase, or laccase, respectively, and an electron transfer agent thatfacilitates the electrooxidation of the glucose, lactate, or oxygen,respectively.

In another embodiment, the sensing layer 64 is not deposited directly onthe working electrode 58 a. Instead, the sensing layer 64 is spacedapart from the working electrode 58 a, as illustrated in FIG. 4A, andseparated from the working electrode 58 a by a separation layer 61. Theseparation layer 61 typically includes one or more membranes or films.In addition to separating the working electrode 58 a from the sensinglayer 64, the separation layer 61 may also act as a mass transportlimiting layer or an interferent eliminating layer, as described below.

Typically, a sensing layer 64, which is not in direct contact with theworking electrode 58 a, includes a catalyst that facilitates a reactionof the analyte. However, this sensing layer 64 typically does notinclude an electron transfer agent that transfers electrons directlyfrom the working electrode 58 a to the analyte, as the sensing layer 64is spaced apart from the working electrode 58 a. One example of thistype of sensor is a glucose or lactate sensor which includes an enzyme(e.g., glucose oxidase or lactate oxidase, respectively) in the sensinglayer 64. The glucose or lactate reacts with a second compound (e.g.,oxygen) in the presence of the enzyme. The second compound is thenelectrooxidized or electroreduced at the electrode. Changes in thesignal at the electrode indicate changes in the level of the secondcompound in the fluid and are proportional to changes in glucose orlactate level and, thus, correlate to the analyte level.

In another embodiment, two sensing layers 63, 64 are used, as shown inFIG. 4B. Each of the two sensing layers 63, 64 may be independentlyformed on the working electrode 58 a or in proximity to the workingelectrode 58 a. One sensing layer 64 is typically, although notnecessarily, spaced apart from the working electrode 58 a. For example,this sensing layer 64 may include a catalyst which catalyzes a reactionof the analyte to form a product compound. The product compound is thenelectrolyzed in the second sensing layer 63 which may include anelectron transfer agent to transfer electrons between the workingelectrode 58 a and the product compound and/or a second catalyst tocatalyze a reaction of the product compound to generate a signal at theworking electrode 58 a.

For example, a glucose or lactate sensor may include a first sensinglayer 64 which is spaced apart from the working electrode and containsan enzyme, for example, glucose oxidase or lactate oxidase. The reactionof glucose or lactate in the presence of the appropriate enzyme formshydrogen peroxide. A second sensing layer 63 is provided directly on theworking electrode 58 a and contains a peroxidase enzyme and an electrontransfer agent to generate a signal at the electrode in response to thehydrogen peroxide. The level of hydrogen peroxide indicated by thesensor then correlates to the level of glucose or lactate. Anothersensor which operates similarly can be made using a single sensing layerwith both the glucose or lactate oxidase and the peroxidase beingdeposited in the single sensing layer. Examples of such sensors aredescribed in U.S. Pat. No. 5,593,852, U.S. Pat. No. 5,665,222, and PCTPatent Application No. US98/02403 entitled “Soybean PeroxidaseElectrochemical Sensor”, filed on Feb. 11, 1998, published asWO-1998/035053, incorporated herein by reference.

In some embodiments, one or more of the working electrodes 58 b do nothave a corresponding sensing layer 64, as shown in FIGS. 3A and 4A, orhave a sensing layer (not shown) which does not contain one or morecomponents (e.g., an electron transfer agent or catalyst) needed toelectrolyze the analyte. The signal generated at this working electrode58 b typically arises from interferents and other sources, such as ions,in the fluid, and not in response to the analyte (because the analyte isnot electrooxidized or electroreduced). Thus, the signal at this workingelectrode 58 b corresponds to a background signal. The background signalcan be removed from the analyte signal obtained from other workingelectrodes 58 a that are associated with fully-functional sensing layers64 by, for example, subtracting the signal at working electrode 58 bfrom the signal at working electrode 58 a.

Sensors having multiple working electrodes 58 a may also be used toobtain more precise results by averaging the signals or measurementsgenerated at these working electrodes 58 a. In addition, multiplereadings at a single working electrode 58 a or at multiple workingelectrodes may be averaged to obtain more precise data.

Electron Transfer Agent

In many embodiments, the sensing layer 64 contains one or more electrontransfer agents in contact with the conductive material 56 of theworking electrode 58, as shown in FIGS. 3A and 3B. In some embodimentsof the invention, there is little or no leaching of the electrontransfer agent away from the working electrode 58 during the period inwhich the sensor 42 is implanted in the patient. A diffusing orleachable (i.e., releasable) electron transfer agent often diffuses intothe analyte-containing fluid, thereby reducing the effectiveness of theelectrode by reducing the sensitivity of the sensor over time. Inaddition, a diffusing or leaching electron transfer agent in animplantable sensor 42 may also cause damage to the patient. In theseembodiments, preferably, at least 90%, more preferably, at least 95%,and, most preferably, at least 99%, of the electron transfer agentremains disposed on the sensor after immersion in the analyte-containingfluid for 24 hours, and, more preferably, for 72 hours. In particular,for an implantable sensor, preferably, at least 90%, more preferably, atleast 95%, and most preferably, at least 99%, of the electron transferagent remains disposed on the sensor after immersion in the body fluidat 37° C. for 24 hours, and, more preferably, for 72 hours.

In some embodiments of the invention, to prevent leaching, the electrontransfer agents are bound or otherwise immobilized on the workingelectrode 58 or between or within one or more membranes or filmsdisposed over the working electrode 58. The electron transfer agent maybe immobilized on the working electrode 58 using, for example, apolymeric or sol-gel immobilization technique. Alternatively, theelectron transfer agent may be chemically (e.g., ionically, covalently,or coordinatively) bound to the working electrode 58, either directly orindirectly through another molecule, such as a polymer, that is in turnbound to the working electrode 58.

Application of the sensing layer 64 on a working electrode 58 a is onemethod for creating a working surface for the working electrode 58 a, asshown in FIGS. 3A and 3B. The electron transfer agent mediates thetransfer of electrons to electrooxidize or electroreduce an analyte andthereby permits a current flow between the working electrode 58 and thecounter electrode 60 via the analyte. The mediation of the electrontransfer agent facilitates the electrochemical analysis of analyteswhich are not suited for direct electrochemical reaction on anelectrode.

In general, the preferred electron transfer agents are electroreducibleand electrooxidizable ions or molecules having redox potentials that area few hundred millivolts above or below the redox potential of thestandard calomel electrode (SCE). Preferably, the electron transferagents are not more reducing than about −150 mV and not more oxidizingthan about +400 mV versus SCE.

The electron transfer agent may be organic, organometallic, orinorganic. Examples of organic redox species are quinones and speciesthat in their oxidized state have quinoid structures, such as Nile blueand indophenol. Some quinones and partially oxidized quinhydrones reactwith functional groups of proteins such as the thiol groups of cysteine,the amine groups of lysine and arginine, and the phenolic groups oftyrosine which may render those redox species unsuitable for some of thesensors of the present invention because of the presence of theinterfering proteins in an analyte-containing fluid. Usually substitutedquinones and molecules with quinoid structure are less reactive withproteins and are preferred. A preferred tetrasubstituted quinone usuallyhas carbon atoms in positions 1, 2, 3, and 4.

In general, electron transfer agents suitable for use in the inventionhave structures or charges which prevent or substantially reduce thediffusional loss of the electron transfer agent during the period oftime that the sample is being analyzed. The preferred electron transferagents include a redox species bound to a polymer which can in turn beimmobilized on the working electrode. The bond between the redox speciesand the polymer may be covalent, coordinative, or ionic. Useful electrontransfer agents and methods for producing them are described in U.S.Pat. Nos. 5,264,104; 5,356,786; 5,262,035; and 5,320,725, incorporatedherein by reference. Although any organic or organometallic redoxspecies can be bound to a polymer and used as an electron transferagent, the preferred redox species is a transition metal compound orcomplex. The preferred transition metal compounds or complexes includeosmium, ruthenium, iron, and cobalt compounds or complexes. The mostpreferred are osmium compounds and complexes. It will be recognized thatmany of the redox species described below may also be used, typicallywithout a polymeric component, as electron transfer agents in a carrierfluid or in a sensing layer of a sensor where leaching of the electrontransfer agent is acceptable.

One type of non-releasable polymeric electron transfer agent contains aredox species covalently bound in a polymeric composition. An example ofthis type of mediator is poly(vinylferrocene).

Another type of non-releasable electron transfer agent contains anionically-bound redox species. Typically, this type of mediator includesa charged polymer coupled to an oppositely charged redox species.Examples of this type of mediator include a negatively charged polymersuch as Nafion® (DuPont) coupled to a positively charged redox speciessuch as an osmium or ruthenium polypyridyl cation. Another example of anionically-bound mediator is a positively charged polymer such asquaternized poly(4-vinyl pyridine) or poly(1-vinyl imidazole) coupled toa negatively charged redox species such as ferricyanide or ferrocyanide.The preferred ionically-bound redox species is a highly charged redoxspecies bound within an oppositely charged redox polymer.

In another embodiment of the invention, suitable non-releasable electrontransfer agents include a redox species coordinatively bound to apolymer. For example, the mediator may be formed by coordination of anosmium or cobalt 2,2′-bipyridyl complex to poly(1-vinyl imidazole) orpoly(4-vinyl pyridine).

The preferred electron transfer agents are osmium transition metalcomplexes with one or more ligands, each ligand having anitrogen-containing heterocycle such as 2,2′-bipyridine,1,10-phenanthroline, or derivatives thereof. Furthermore, the preferredelectron transfer agents also have one or more ligands covalently boundin a polymer, each ligand having at least one nitrogen-containingheterocycle, such as pyridine, imidazole, or derivatives thereof. Thesepreferred electron transfer agents exchange electrons rapidly betweeneach other and the working electrodes 58 so that the complex can berapidly oxidized and reduced.

One example of a particularly useful electron transfer agent includes(a) a polymer or copolymer having pyridine or imidazole functionalgroups and (b) osmium cations complexed with two ligands, each ligandcontaining 2,2′-bipyridine, 1,10-phenanthroline, or derivatives thereof,the two ligands not necessarily being the same. Preferred derivatives of2,2′-bipyridine for complexation with the osmium cation are4,4′-dimethyl-2,2′-bipyridine and mono-, di-, andpolyalkoxy-2,2′-bipyridines, such as 4,4′-dimethoxy-2,2′-bipyridine.Preferred derivatives of 1,10-phenanthroline for complexation with theosmium cation are 4,7-dimethyl-1,10-phenanthroline and mono, di-, andpolyalkoxy-1,10-phenanthrolines, such as4,7-dimethoxy-1,10-phenanthroline. Preferred polymers for complexationwith the osmium cation include polymers and copolymers of poly(1-vinylimidazole) (referred to as “PVI”) and poly(4-vinyl pyridine) (referredto as “PVP”). Suitable copolymer substituents of poly(1-vinyl imidazole)include acrylonitrile, acrylamide, and substituted or quaternizedN-vinyl imidazole. Most preferred are electron transfer agents withosmium complexed to a polymer or copolymer of poly(1-vinyl imidazole).

The preferred electron transfer agents have a redox potential rangingfrom −100 mV to about +150 mV versus the standard calomel electrode(SCE). Preferably, the potential of the electron transfer agent rangesfrom −100 mV to +150 mV and more preferably, the potential ranges from−50 mV to +50 mV. The most preferred electron transfer agents haveosmium redox centers and a redox potential ranging from +50 mV to −150mV versus SCE.

Catalyst

The sensing layer 64 may also include a catalyst which is capable ofcatalyzing a reaction of the analyte. The catalyst may also, in someembodiments, act as an electron transfer agent. One example of asuitable catalyst is an enzyme which catalyzes a reaction of theanalyte. For example, a catalyst, such as a glucose oxidase, glucosedehydrogenase (e.g., pyrroloquinoline quinone glucose dehydrogenase(PQQ)), or oligosaccharide dehydrogenase, may be used when the analyteis glucose. A lactate oxidase or lactate dehydrogenase may be used whenthe analyte is lactate. Laccase may be used when the analyte is oxygenor when oxygen is generated or consumed in response to a reaction of theanalyte.

Preferably, the catalyst is non-leachably disposed on the sensor,whether the catalyst is part of a solid sensing layer in the sensor orsolvated in a fluid within the sensing layer. More preferably, thecatalyst is immobilized within the sensor (e.g., on the electrode and/orwithin or between a membrane or film) to prevent unwanted leaching ofthe catalyst away from the working electrode 58 and into the patient.This may be accomplished, for example, by attaching the catalyst to apolymer, cross linking the catalyst with another electron transfer agent(which, as described above, can be polymeric), and/or providing one ormore barrier membranes or films with pore sizes smaller than thecatalyst.

As described above, a second catalyst may also be used. This secondcatalyst is often used to catalyze a reaction of a product compoundresulting from the catalyzed reaction of the analyte. The secondcatalyst typically operates with an electron transfer agent toelectrolyze the product compound to generate a signal at the workingelectrode. Alternatively, the second catalyst may be provided in aninterferent-eliminating layer to catalyze reactions that removeinterferents, as described below.

One embodiment of the invention is an electrochemical sensor in whichthe catalyst is mixed or dispersed in the conductive material 56 whichforms the conductive trace 52 of a working electrode 58. This may beaccomplished, for example, by mixing a catalyst, such as an enzyme, in acarbon ink and applying the mixture into a channel 54 on the surface ofthe substrate 50. Preferably, the catalyst is immobilized in the channel53 so that it can not leach away from the working electrode 58. This maybe accomplished, for example, by curing a binder in the carbon ink usinga curing technique appropriate to the binder. Curing techniques include,for example, evaporation of a solvent or dispersant, exposure toultraviolet light, or exposure to heat. Typically, the mixture isapplied under conditions that do not substantially degrade the catalyst.For example, the catalyst may be an enzyme that is heat-sensitive. Theenzyme and conductive material mixture should be applied and cured,preferably, without sustained periods of heating. The mixture may becured using evaporation or UV curing techniques or by the exposure toheat that is sufficiently short that the catalyst is not substantiallydegraded.

Another consideration for in vivo analyte sensors is the thermostabilityof the catalyst. Many enzymes have only limited stability at biologicaltemperatures. Thus, it may be necessary to use large amounts of thecatalyst and/or use a catalyst that is thermostable at the necessarytemperature (e.g., 37° C. or higher for normal body temperature). Athermostable catalyst may be defined as a catalyst which loses less than5% of its activity when held at 37° C. for at least one hour,preferably, at least one day, and more preferably at least three days.One example of a thermostable catalyst is soybean peroxidase. Thisparticular thermostable catalyst may be used in a glucose or lactatesensor when combined either in the same or separate sensing layers withglucose or lactate oxidase or dehydrogenase. A further description ofthermostable catalysts and their use in electrochemical inventions isfound in U.S. Pat. No. 5,665,222, and PCT Application No. US98/02403entitled “Soybean Peroxidase Electrochemical Sensor”, filed on Feb. 11,1998, published as WO-1998/035053, incorporated herein by reference.

Electrolysis of the Analyte

To electrolyze the analyte, a potential (versus a reference potential)is applied across the working and counter electrodes 58, 60. The minimummagnitude of the applied potential is often dependent on the particularelectron transfer agent, analyte (if the analyte is directlyelectrolyzed at the electrode), or second compound (if a secondcompound, such as oxygen or hydrogen peroxide, whose level is dependenton the analyte level, is directly electrolyzed at the electrode). Theapplied potential usually equals or is more oxidizing or reducing,depending on the desired electrochemical reaction, than the redoxpotential of the electron transfer agent, analyte, or second compound,whichever is directly electrolyzed at the electrode. The potential atthe working electrode is typically large enough to drive theelectrochemical reaction to or near completion.

The magnitude of the potential may optionally be limited to preventsignificant (as determined by the current generated in response to theanalyte) electrochemical reaction of interferents, such as urate,ascorbate, and acetaminophen. The limitation of the potential may beobviated if these interferents have been removed in another way, such asby providing an interferent-limiting barrier, as described below, or byincluding a working electrode 58 b (see FIG. 3A) from which a backgroundsignal may be obtained.

When a potential is applied between the working electrode 58 and thecounter electrode 60, an electrical current will flow. The current is aresult of the electrolysis of the analyte or a second compound whoselevel is affected by the analyte. In one embodiment, the electrochemicalreaction occurs via an electron transfer agent and the optionalcatalyst. Many analytes B are oxidized (or reduced) to products C by anelectron transfer agent species A in the presence of an appropriatecatalyst (e.g., an enzyme). The electron transfer agent A is thenoxidized (or reduced) at the electrode. Electrons are collected by (orremoved from) the electrode and the resulting current is measured. Thisprocess is illustrated by reaction equations (1) and (2) (similarequations may be written for the reduction of the analyte B by a redoxmediator A in the presence of a catalyst):

$\begin{matrix}{{{{nA}({ox})} + B}\overset{catalyst}{\rightarrow}{{{nA}({red})} + C}} & (1) \\{{{{nA}({red})}\overset{electrode}{\rightarrow}{{{nA}({ox})} + {ne}}}} & (2)\end{matrix}$

As an example, an electrochemical sensor may be based on the reaction ofa glucose molecule with two non-leachable ferricyanide anions in thepresence of glucose oxidase to produce two non-leachable ferrocyanideanions, two hydrogen ions, and gluconolactone. The amount of glucosepresent is assayed by electrooxidizing the non-leachable ferrocyanideanions to non-leachable ferricyanide anions and measuring the current.

In another embodiment, a second compound whose level is affected by theanalyte is electrolyzed at the working electrode. In some cases, theanalyte D and the second compound, in this case, a reactant compound E,such as oxygen, react in the presence of the catalyst, as shown inreaction equation (3).

$\begin{matrix}{{D + E}\overset{catalyst}{\rightarrow}{F + G}} & (3)\end{matrix}$

The reactant compound E is then directly oxidized (or reduced) at theworking electrode, as shown in reaction equation (4)

$\begin{matrix}{{{nE}({red})}\overset{electrode}{\rightarrow}{{{nE}({ox})} + {ne}}} & (4)\end{matrix}$

Alternatively, the reactant compound E is indirectly oxidized (orreduced) using an electron transfer agent H (optionally in the presenceof a catalyst), that is subsequently reduced or oxidized at theelectrode, as shown in reaction equations (5) and (6).

$\begin{matrix}{{{{nH}({ox})} + E}->{{{nH}({red})} + I}} & (5) \\{{{nH}({red})}\overset{electrode}{\rightarrow}{{{nH}({ox})} + {ne}}} & (6)\end{matrix}$

In either case, changes in the concentration of the reactant compound,as indicated by the signal at the working electrode, correspondinversely to changes in the analyte (i.e., as the level of analyteincrease then the level of reactant compound and the signal at theelectrode decreases.)

In other embodiments, the relevant second compound is a product compoundF, as shown in reaction equation (3). The product compound F is formedby the catalyzed reaction of analyte D and then be directly electrolyzedat the electrode or indirectly electrolyzed using an electron transferagent and, optionally, a catalyst. In these embodiments, the signalarising from the direct or indirect electrolysis of the product compoundF at the working electrode corresponds directly to the level of theanalyte (unless there are other sources of the product compound). As thelevel of analyte increases, the level of the product compound and signalat the working electrode increases.

Those skilled in the art will recognize that there are many differentreactions that will achieve the same result; namely the electrolysis ofan analyte or a compound whose level depends on the level of theanalyte. Reaction equations (1) through (6) illustrate non-limitingexamples of such reactions.

Temperature Probe

A variety of optional items may be included in the sensor. One optionalitem is a temperature probe 66 (FIGS. 8 and 11). The temperature probe66 may be made using a variety of known designs and materials. Oneexemplary temperature probe 66 is formed using two probe leads 68, 70connected to each other through a temperature-dependent element 72 thatis formed using a material with a temperature-dependent characteristic.An example of a suitable temperature-dependent characteristic is theresistance of the temperature-dependent element 72.

The two probe leads 68, 70 are typically formed using a metal, an alloy,a semimetal, such as graphite, a degenerate or highly dopedsemiconductor, or a small-band gap semiconductor. Examples of suitablematerials include gold, silver, ruthenium oxide, titanium nitride,titanium dioxide, indium doped tin oxide, tin doped indium oxide, orgraphite. The temperature-dependent element 72 is typically made using afine trace (e.g., a conductive trace that has a smaller cross-sectionthan that of the probe leads 68, 70) of the same conductive material asthe probe leads, or another material such as a carbon ink, a carbonfiber, or platinum, which has a temperature-dependent characteristic,such as resistance, that provides a temperature-dependent signal when avoltage source is attached to the two probe leads 68, 70 of thetemperature probe 66. The temperature-dependent characteristic of thetemperature-dependent element 72 may either increase or decrease withtemperature. Preferably, the temperature dependence of thecharacteristic of the temperature-dependent element 72 is approximatelylinear with temperature over the expected range of biologicaltemperatures (about 25 to 45° C.), although this is not required.

Typically, a signal (e.g., a current) having an amplitude or otherproperty that is a function of the temperature can be obtained byproviding a potential across the two probe leads 68, 70 of thetemperature probe 66. As the temperature changes, thetemperature-dependent characteristic of the temperature-dependentelement 72 increases or decreases with a corresponding change in thesignal amplitude. The signal from the temperature probe 66 (e.g., theamount of current flowing through the probe) may be combined with thesignal obtained from the working electrode 58 by, for example, scalingthe temperature probe signal and then adding or subtracting the scaledtemperature probe signal from the signal at the working electrode 58. Inthis manner, the temperature probe 66 can provide a temperatureadjustment for the output from the working electrode 58 to offset thetemperature dependence of the working electrode 58.

One embodiment of the temperature probe includes probe leads 68, 70formed as two spaced-apart channels with a temperature-dependent element72 formed as a cross-channel connecting the two spaced-apart channels,as illustrated in FIG. 8. The two spaced-apart channels contain aconductive material, such as a metal, alloy, semimetal, degeneratesemiconductor, or metallic compound. The cross-channel may contain thesame material (provided the cross-channel has a smaller cross-sectionthan the two spaced-apart channels) as the probe leads 68, 70. In otherembodiments, the material in the cross-channel is different than thematerial of the probe leads 68, 70.

One exemplary method for forming this particular temperature probeincludes forming the two spaced-apart channels and then filling themwith the metallic or alloyed conductive material. Next, thecross-channel is formed and then filled with the desired material. Thematerial in the cross-channel overlaps with the conductive material ineach of the two spaced-apart channels to form an electrical connection.

For proper operation of the temperature probe 66, thetemperature-dependent element 72 of the temperature probe 66 can not beshorted by conductive material formed between the two probe leads 68,70. In addition, to prevent conduction between the two probe leads 68,70 by ionic species within the body or sample fluid, a covering may beprovided over the temperature-dependent element 72, and preferably overthe portion of the probe leads 68, 70 that is implanted in the patient.The covering may be, for example, a non-conducting film disposed overthe temperature-dependent element 72 and probe leads 68, 70 to preventthe ionic conduction. Suitable non-conducting films include, forexample, Kapton™ polyimide films (DuPont, Wilmington, Del.).

Another method for eliminating or reducing conduction by ionic speciesin the body or sample fluid is to use an ac voltage source connected tothe probe leads 68, 70. In this way, the positive and negative ionicspecies are alternately attracted and repelled during each half cycle ofthe ac voltage. This results in no net attraction of the ions in thebody or sample fluid to the temperature probe 66. The maximum amplitudeof the ac current through the temperature-dependent element 72 may thenbe used to correct the measurements from the working electrodes 58.

The temperature probe can be placed on the same substrate as theelectrodes. Alternatively, a temperature probe may be placed on aseparate substrate. In addition, the temperature probe may be used byitself or in conjunction with other devices.

Biocompatible Layer

An optional film layer 75 is formed over at least that portion of thesensor 42 which is subcutaneously inserted into the patient, as shown inFIG. 9. This optional film layer 74 may serve one or more functions. Thefilm layer 74 prevents the penetration of large biomolecules into theelectrodes. This is accomplished by using a film layer 74 having a poresize that is smaller than the biomolecules that are to be excluded. Suchbiomolecules may foul the electrodes and/or the sensing layer 64 therebyreducing the effectiveness of the sensor 42 and altering the expectedsignal amplitude for a given analyte concentration. The fouling of theworking electrodes 58 may also decrease the effective life of the sensor42. The biocompatible layer 74 may also prevent protein adhesion to thesensor 42, formation of blood clots, and other undesirable interactionsbetween the sensor 42 and body.

For example, the sensor may be completely or partially coated on itsexterior with a biocompatible coating. A preferred biocompatible coatingis a hydrogel which contains at least 20 wt. % fluid when in equilibriumwith the analyte-containing fluid. Examples of suitable hydrogels aredescribed in U.S. Pat. No. 5,593,852, incorporated herein by reference,and include crosslinked polyethylene oxides, such as polyethylene oxidetetraacrylate.

Interferent-Eliminating Layer

An interferent-eliminating layer (not shown) may be included in thesensor 42. The interferent-eliminating layer may be incorporated in thebiocompatible layer 75 or in the mass transport limiting layer 74(described below) or may be a separate layer. Interferents are moleculesor other species that are electroreduced or electrooxidized at theelectrode, either directly or via an electron transfer agent, to producea false signal. In one embodiment, a film or membrane prevents thepenetration of one or more interferents into the region around theworking electrodes 58. Preferably, this type of interferent-eliminatinglayer is much less permeable to one or more of the interferents than tothe analyte.

The interferent-eliminating layer may include ionic components, such asNafion®, incorporated into a polymeric matrix to reduce the permeabilityof the interferent-eliminating layer to ionic interferents having thesame charge as the ionic components. For example, negatively chargedcompounds or compounds that form negative ions may be incorporated inthe interferent-eliminating layer to reduce the permeation of negativespecies in the body or sample fluid.

Another example of an interferent-eliminating layer includes a catalystfor catalyzing a reaction which removes interferents. One example ofsuch a catalyst is a peroxidase. Hydrogen peroxide reacts withinterferents, such as acetaminophen, urate, and ascorbate. The hydrogenperoxide may be added to the analyte-containing fluid or may begenerated in situ, by, for example, the reaction of glucose or lactatein the presence of glucose oxidase or lactate oxidase, respectively.Examples of interferent eliminating layers include a peroxidase enzymecrosslinked (a) using gluteraldehyde as a crosslinking agent or (b)oxidation of oligosaccharide groups in the peroxidase glycoenzyme withNaIO4, followed by coupling of the aldehydes formed to hydrazide groupsin a polyacrylamide matrix to form hydrazones are describe in U.S. Pat.Nos. 5,262,305 and 5,356,786, incorporated herein by reference.

Mass Transport Limiting Layer

A mass transport limiting layer 74 may be included with the sensor toact as a diffusion-limiting barrier to reduce the rate of mass transportof the analyte, for example, glucose or lactate, into the region aroundthe working electrodes 58. By limiting the diffusion of the analyte, thesteady state concentration of the analyte in the proximity of theworking electrode 58 (which is proportional to the concentration of theanalyte in the body or sample fluid) can be reduced. This extends theupper range of analyte concentrations that can still be accuratelymeasured and may also expand the range in which the current increasesapproximately linearly with the level of the analyte.

It is preferred that the permeability of the analyte through the filmlayer 74 vary little or not at all with temperature, so as to reduce oreliminate the variation of current with temperature. For this reason, itis preferred that in the biologically relevant temperature range fromabout 25° C. to about 45° C., and most importantly from 30° C. to 40°C., neither the size of the pores in the film nor its hydration orswelling change excessively. Preferably, the mass transport limitinglayer is made using a film that absorbs less than 5 wt. % of fluid over24 hours. This may reduce or obviate any need for a temperature probe.For implantable sensors, it is preferable that the mass transportlimiting layer is made using a film that absorbs less than 5 wt. % offluid over 24 hours at 37° C.

Particularly useful materials for the film layer 74 are membranes thatdo not swell in the analyte-containing fluid that the sensor tests.Suitable membranes include 3 to 20,000 nm diameter pores. Membraneshaving 5 to 500 nm diameter pores with well-defined, uniform pore sizesand high aspect ratios are preferred. In one embodiment, the aspectratio of the pores is preferably two or greater and more preferably fiveor greater.

Well-defined and uniform pores can be made by track etching a polymericmembrane using accelerated electrons, ions, or particles emitted byradioactive nuclei. Most preferred are anisotropic, polymeric, tracketched membranes that expand less in the direction perpendicular to thepores than in the direction of the pores when heated. Suitable polymericmembranes included polycarbonate membranes from Poretics (Livermore,Calif., catalog number 19401, 0.01 μm pore size polycarbonate membrane)and Corning Costar Corp. (Cambridge, Mass., Nucleopore™ brand membraneswith 0.015 μm pore size). Other polyolefin and polyester films may beused. It is preferred that the permeability of the mass transportlimiting membrane changes no more than 4%, preferably, no more than 3%,and, more preferably, no more than 2%, per ° C. in the range from 30° C.to 40° C. when the membranes resides in the subcutaneous interstitialfluid.

In some embodiments of the invention, the mass transport limiting layer74 may also limit the flow of oxygen into the sensor 42. This canimprove the stability of sensors 42 that are used in situations wherevariation in the partial pressure of oxygen causes non-linearity insensor response. In these embodiments, the mass transport limiting layer74 restricts oxygen transport by at least 40%, preferably at least 60%,and more preferably at least 80%, than the membrane restricts transportof the analyte. For a given type of polymer, films having a greaterdensity (e.g., a density closer to that of the crystalline polymer) arepreferred. Polyesters, such as polyethylene terephthalate, are typicallyless permeable to oxygen and are, therefore, preferred overpolycarbonate membranes.

Anticlotting Agent

An implantable sensor may also, optionally, have an anticlotting agentdisposed on a portion the substrate which is implanted into a patient.This anticlotting agent may reduce or eliminate the clotting of blood orother body fluid around the sensor, particularly after insertion of thesensor. Blood clots may foul the sensor or irreproducibly reduce theamount of analyte which diffuses into the sensor. Examples of usefulanticlotting agents include heparin and tissue plasminogen activator(TPA), as well as other known anticlotting agents.

The anticlotting agent may be applied to at least a portion of that partof the sensor 42 that is to be implanted. The anticlotting agent may beapplied, for example, by bath, spraying, brushing, or dipping. Theanticlotting agent is allowed to dry on the sensor 42. The anticlottingagent may be immobilized on the surface of the sensor or it may beallowed to diffuse away from the sensor surface. Typically, thequantities of anticlotting agent disposed on the sensor are far belowthe amounts typically used for treatment of medical conditions involvingblood clots and, therefore, have only a limited, localized effect.

Sensor Lifetime

The sensor 42 may be designed to be a replaceable component in an invivo or in vitro analyte monitor, and particularly in an implantableanalyte monitor. Typically, the sensor 42 is capable of operation over aperiod of days. Preferably, the period of operation is at least one day,more preferably at least three days, and most preferably at least oneweek. The sensor 42 can then be removed and replaced with a new sensor.The lifetime of the sensor 42 may be reduced by the fouling of theelectrodes or by the leaching of the electron transfer agent orcatalyst. These limitations on the longevity of the sensor 42 can beovercome by the use of a biocompatible layer 75 or non-leachableelectron transfer agent and catalyst, respectively, as described above.

Another primary limitation on the lifetime of the sensor 42 is thetemperature stability of the catalyst. Many catalysts are enzymes, whichare very sensitive to the ambient temperature and may degrade attemperatures of the patient's body (e.g., approximately 37° C. for thehuman body). Thus, robust enzymes should be used where available. Thesensor 42 should be replaced when a sufficient amount of the enzyme hasbeen deactivated to introduce an unacceptable amount of error in themeasurements.

The present invention should not be considered limited to the particularexamples described above, but rather should be understood to cover allaspects of the invention as fairly set out in the attached claims.Various modifications, equivalent processes, as well as numerousstructures to which the present invention may be applicable will bereadily apparent to those of skill in the art to which the presentinvention is directed upon review of the instant specification. Theclaims are intended to cover such modifications and devices.

1. (canceled)
 2. An analyte sensor for measuring an analyte in a host, the sensor comprising: a first working electrode disposed beneath an active enzymatic portion of a membrane; and a second working electrode disposed beneath an inactive-enzymatic portion of a membrane, wherein the inactive enzymatic portion comprises at least one of a deactivated enzyme or an inactive enzyme.
 3. The sensor of claim 2, further comprising a reference electrode.
 4. The sensor of claim 3, further comprising a counter electrode.
 5. The sensor of claim 2, wherein the membrane located over the first working electrode and the membrane located over the second working electrode each comprise an interference domain that restricts a flow of at least one interfering species.
 6. The sensor of claim 5, wherein the interference domain comprises a material selected from the group consisting of a polyurethane and a cellulosic polymer.
 7. The sensor of claim 2, wherein the membrane located over the first working electrode and the membrane located over the second working electrode each comprise a resistance domain that controls a flux of an analyte therethrough.
 8. The sensor of claim 7, wherein the resistance domain comprises a material selected from the group consisting of a polyurethane and a silicone.
 9. The sensor of claim 2, wherein the analyte sensor is a glucose sensor, and wherein the first working electrode is configured to generate a first signal associated with glucose related electroactive compounds and non-glucose related electroactive compounds, wherein the glucose related electroactive compounds and the non-glucose related electroactive compounds have a first oxidation potential.
 10. The sensor of claim 9, wherein the second working electrode is configured to generate a second signal associated with non-glucose related electroactive compounds, wherein the non-glucose related electroactive compounds have an oxidation potential that substantially overlaps with the first oxidation potential.
 11. An analyte sensor for measuring an analyte in a host, the sensor comprising: a first working electrode disposed beneath an active enzymatic portion of a membrane; a second working electrode disposed beneath an inactive-enzymatic or non-enzymatic portion of a membrane; and an insulator located between the first working electrode and the second working electrode, wherein the first working electrode and the second working electrode are intertwined.
 12. An analyte sensor for measuring an analyte concentration in a host, the sensor comprising: a first working electrode configured to generate a first signal associated with analyte related electroactive compounds and non-analyte related electroactive compounds, wherein the analyte related electroactive compounds and the non-analyte related electroactive compounds have a first oxidation potential; and a second working electrode configured to generate a second signal associated with non-analyte related electroactive compounds, wherein the non-analyte related electroactive compounds have an oxidation potential that substantially overlaps with the first oxidation potential.
 13. An analyte sensor for measuring an analyte concentration in a host, the sensor comprising: a first working electrode configured to generate a first signal associated with analyte related electroactive compounds and non-analyte related electroactive compounds, wherein the analyte related electroactive compounds and the non-analyte related electroactive compounds have a first oxidation potential; and a second working electrode configured to generate a second signal associated with non-analyte related electroactive compounds.
 14. The sensor of claim 13, wherein the non-analyte related electroactive compounds have an oxidation potential that substantially overlaps with the first oxidation potential.
 15. An analyte sensor for measuring an analyte in a host, the sensor comprising: a first working electrode comprising a first electroactive surface disposed beneath an active enzymatic portion of a membrane, wherein the first working electrode defines a first longitudinal axis; a second working electrode comprising a second electroactive surface disposed beneath an inactive-enzymatic portion of a membrane or a non-enzymatic portion of a membrane; and a flow path diffusion barrier configured to substantially block diffusion of at least one of an analyte and a co-analyte between the first electroactive surface and the second electroactive surface by an offset of the first electroactive surface and the second electroactive surface along a longitudinal axis of the sensor.
 16. The sensor of claim 15, wherein the second working electrode defines a second longitudinal axis that is co-linear with the first longitudinal axis, wherein both the first longitudinal axis and the second longitudinal axis define the longitudinal axis of the sensor.
 17. The sensor of claim 15, wherein the sensor is configured for contact with a blood flow from a circulatory system of a host.
 18. An analyte sensor for measuring an analyte in a host, the sensor comprising: a first working electrode comprising a first electroactive surface disposed beneath an active enzymatic portion of a membrane, wherein the first working electrode defines a first longitudinal axis; a counter electrode comprising a second electroactive surface disposed beneath an inactive-enzymatic portion of a membrane or a non-enzymatic portion of a membrane, wherein the second working electrode defines a second longitudinal axis that is co-linear with the first longitudinal axis, wherein both the first longitudinal axis and the second longitudinal axis define a longitudinal axis of the sensor; and a flow path diffusion barrier configured to substantially block diffusion of at least one of an analyte and a co-analyte between the first electroactive surface and the second electroactive surface by an offset of the first electroactive surface and the second electroactive surface along the longitudinal axis of the sensor.
 19. The sensor of claim 18, wherein the sensor is configured for contact with a blood flow from a circulatory system of a host. 